Magnetic resonance imaging (MRI) is a technique that is capable of providing three-dimensional imaging of an object. A conventional MRI system typically includes a main or primary magnet that provides the main static magnetic field Bo, magnetic field gradient coils and radio frequency (RF) coils, which are used for spatial encoding, exciting and detecting the nuclei for imaging. Typically, the main magnet is designed to provide a homogeneous magnetic field in an internal region within the main magnet, for example, in the air space of a large central bore of a solenoid or in the air gap between the magnetic pole plates of a C-type magnet. The patient or object to be imaged is positioned in the homogeneous field region located in such air space. The gradient field and the RF coils are typically located external to the patient or object to be imaged and inside the geometry of the main or primary magnet(s) surrounding the air space. There is shown in U.S. Pat. Nos. 4,689,563; 4,968,937 and 5,990,681, the teachings of which are incorporated herein by reference, some exemplary MRI systems.
In MRI, the uniform magnetic field Bo generated by the main magnet is applied to an imaged object by convention along the Z-axis of a Cartesian coordinate system, the origin of which is within the imaged object. The uniform magnetic field Bo being applied has the effect of aligning the magnetization arising from the nuclei of the atoms comprising the imaged object, along the Z-axis, such nuclei possess a nuclear magnetization due to their having an odd number of protons or neutrons. In response to RF magnetic field pulses of the proper frequency, with field direction orientated within the XY plane, the nuclei resonate at their Larmor frequencies, ω=γBo, where γ is called the gyromagnetic ratio. In a typical planar imaging sequence, the RF signal centered about the desired Larmor frequency is applied to the imaged object at the same time a magnetic field gradient Gz is being applied along the Z-axis. This gradient field Gz causes only the nuclei in a slice of limited thickness through the object perpendicular to the Z-axis, to satisfy the resonant condition and thus be excited into resonance.
After excitation of the nuclei in the slice, magnetic field gradients are applied along the X- and Y-axes respectively. The gradient Gx along the X-axis causes the nuclei to precess at different frequencies depending on their position along the X-axis, that is, Gx spatially encodes the processing nuclei by frequency. Thus, this gradient is often referred to as a frequency encoding or read-out gradient. The Y-axis gradient Gy is incremented through a series of values and encodes the Y position into the rate of change of the phase of the precessing nuclei as a function of gradient amplitude, a process typically referred to as phase encoding.
The quality of the image produced by the MRI techniques is dependent, in part, upon the strength of the magnetic resonance (MR) signal received from the precessing nuclei by the MRI detector. Loop surface coil detectors were adopted in the early 1980's for MRI from prior uses in NMR. There is described, for example, in U.S. Pat. No. 4,825,162 a surface coil(s) for use in MRI/NMRI imaging and methods related thereto. In the preferred embodiment of that invention, each surface coil is connected to the input of an associated one of a like plurality of low-input-impedance preamplifiers, which minimize the interaction between any surface coil and any other surface coils not immediately adjacent thereto. These surface coils can have square, circular and the like geometries. This yields an array of a plurality of closely spaced surface coils, each positioned so as to have substantially no interaction with all adjacent surface coils. A different MR response signal is received at each different one of the surface coils from an associated portion of the sample enclosed within the imaging volume defined by the array. Each different MR response signal is used to construct a different one of plurality of different images from each surface coil. These images are then being combined, on a point-by-point basis to produce a single composite MR image of a total sample portion comprised of the MR response signals from the entire array of surface coils.
The use of a phased array of RF coils or surface coils with a tuned and matched circuit including low impedance pre-amplifiers have been used to de-couple adjacent loops as a mechanism for improving the signal-to-noise ratio (SNR) and field of view (FOV). In this regard, it should be understood that the term “coupling” refers to the coupling of a signal (e.g., MR signal and/or noise signal and/or RF excite signal) in one coil to an adjacent coil(s), such that the signal being outputted by the adjacent coil is a combination of the signal detected by the adjacent coil and the coupled signal. Consequently, the image from the adjacent coil may be distorted or the SNR degraded to some degree. Although overlapping adjacent coil(s) and using low impedance pre-amplifiers have been effective in minimizing decoupling artifacts and SNR degradation, such circuitry becomes less effective as the number of coils and/or the coil density is increased. In particular, as the spacing between adjacent coils and between adjacent portions of a coil is decreased signal coupling effects increase and become less manageable by the various measures deployed to counter such effects.
There is found in U.S. Pat. No. 6,771,070 a new type of phased-array detector, a planar strip array (PSA), recently introduced for conventional and parallel spatially-encoded MRI. Instead of an array of conventional loop coils, the PSA consists of parallel conducting strips covered and separated from a virtual conducting ground plane by a low-loss dielectric substrate. Each strip serves as an individual detector and is connected to its own preamplifier. The strips are tuned by adjusting their lengths, L, and separation, h, from the ground plane to λ/4, λ/2, or multiples thereof, where λ is the electromagnetic (EM) wavelength at the MRI frequency. The original PSA exhibits important advantages over conventional loop-coil phased arrays. The mutual coupling between detectors is essentially eliminated, and the self-resonance frequencies of strips are intrinsically higher than loop detectors and can potentially outperform loops at frequencies where they are no longer tunable. This suits the PSA well for large arrays, massively parallel MRI, and high-frequency MRI [Zhang X, Ugurbil K, Chen W. Microstrip RF surface coil design for extremely high-field MRI and spectroscopy. Magn Reson Med 2001; 46: 443-450]. Because of the limited ability to adjust the substrate's dielectric properties in the original PSA, strip length is effectively fixed by tuning, and is therefore not a design parameter that can be flexibly adjusted to optimize the signal-to-noise ratio (SNR) performance for specific applications.
There is found in U.S. Pat. No. 7,088,104, a phased-array detector that includes a strip array antenna including a plurality or more of conductors and a plurality of reactive tuning components, where at least one of the plurality of reactive components is electrically coupled to each conductor as well as to ground/virtual ground. The apparent electrical length of the conductors is tuned with the reactive tuning components so the apparent electrical length of the conductor is equal to be about nλ/4, where n is an integer ≧1 and λ is the wavelength of the signal to be detected. In sum, n distributed lumped tuning elements were introduced to permit the λ/4 or λ/2 conditions to be met with strip lengths that can be selected based on the performance requirements of the specific application.
The intrinsic SNR (ISNR), defined as the SNR that a detector could achieve if the detector and MRI system losses are excluded [Edelstein W A, Glover G H, Hardy C J, Redington R W. The intrinsic signal-to noise ratio in NMR imaging. Magn Resort Med 1986; 3: 604-618], measures the potential SNR performance of a particular detector geometry, and therefore permits performance comparisons of different detector geometries. For example, in the quasistatic field limit, the geometry of the loop detector that produces the maximum ISNR has been determined both numerically [Roemer P B, Edelstein W A, Ultimate Sensitivity Limits of Surface Coils. Proc SMRM 6th Annual Meeting 1987; p. 410], and analytically [Chen C N, Hoult D I. Biomedical Magnetic Resonance Technology. Adam Hilger, 1989; 160-161]. Thus, for a region of interest lying at depth d in a semi-infinite planar sample of the sample portion of interest the loop geometry delivering the maximum ISNR has radiusa=d/√/5  {Eq. 1}
This relationship is central to optimizing SNR gains from head and body phased-arrays fabricated from individual loop coil elements. The MRI strip detector differs fundamentally from a conventional loop detector with respect to its geometry, field distribution, sensitivity profile, and that a larger portion of its energy, if activated, may be stored as E-field within the dielectric substrate close to the strip, compared to energy stored as B-field close to the loop. Accordingly, the strip detector exhibits spatial sensitivity and ISNR characteristics that are entirely different to loop detectors. Thus, the usual design rules for optimizing loop detectors are not applicable to MRI detectors embodying one or more strips.
Quadrature MRI detectors are comprised of elements whose principal RF magnetic field components are orthogonal or nearly-orthogonal, such that they can directly receive the circularly polarized MRI or MRS signals from a sample [Hoult, D. I., C. N. Chen, and V. J. Sank, Quadrature detection in the laboratory frame. Magn Reson Med, 1984. 1(3): p. 339-53; Chen, C., Hoult D I, Sank V J, Quadrature Detection Coils—A Further sqrt(2) Improvement in Sensitivity. Journal of Magnetic Resonance, 1983. 54: p. 324-327; Hyde, J. S., et al., Quadrature detection surface coil. Magn Reson Med, 1987. 4(2): p. 179-84; Bottomley, P. A., et al., Proton-decoupled, Overhauser-enhanced, spatially localized carbon-13 spectroscopy in humans. Magn Reson Med, 1989. 12(3): p. 348-63]. By combining the raw signals from a pair of quadrature elements with a hybrid that provides the appropriate 90° phase shift, or by means of a root-of-the-sum-of-the-squares combination of the processed signals, a quadrature detector coil can provide up to a √2-fold gain in the signal-to-noise-ratio (SNR) compared to coils of the same geometry used as linear detectors [Hoult, D. I., C. N. Chen, and V. J. Sank, Quadrature detection in the laboratory frame. Magn Reson Med, 1984. 1(3): p. 339-53; Redpath, T. W., Quadrature rf coil pairs. Magn Reson Med, 1986. 3(1): p. 118-9].
Quadrature surface coils comprised of a circular loop and a ‘figure-8’ or ‘butterfly’ coil with long axis perpendicular to the main magnetic field, B0, have been adopted for both multi-nuclear MRS, and for commercial MRI phased-arrays in applications including the spine where their use with quadrature hybrids halves the number of MRI receiver channels required to service the array.
It thus, would be desirable to provide rules particularly suitable for a single tuned strip detector element of a tunable PSA as well as methods embodying the use of such rules. It would be desirable to provide detection devices that embody such rules as well as apparatuses and systems that embody such detection devices. It also would be desirable to provide quadrature hybrid detection devices that embody a surface coil and phased strip array antenna.